ES2375724T3 - MICROFLUDE DEVICE FOR SEPERATION OF CELLS AND ITS USES. - Google Patents

Abstract

A microfluidic device comprising: (a) a first region of fixed obstacles arranged in a microfluidic channel defining a liquid flow path, in which the obstacles of the first region preferably bind to a first type of cell compared to a second type of cell, in which the obstacles are placed in at least two columns and at least two rows, in which the rows are placed perpendicular to the liquid flow path in relation to the obstacles in the previous row, thus forming a set of equilateral triangular obstacles; and (b) a second region of fixed obstacles arranged in the microfluidic channel, in which the obstacles of the second region are preferably joined to a third type of cell compared to a fourth type of cell, in which the obstacles are placed in at least two columns and at least two rows, in which the rows are placed perpendicular to the liquid flow path and the obstacles of each successive row are displaced in a direction perpendicular to the liquid flow path in relation to the obstacles of the previous row, thus forming a set of triangular equilateral obstacles, in which the second region is located beyond the first region in the microfluidic channel.

Description

Microfluidic device for cell separation and its uses

Background of the invention

The invention relates to the fields of medical diagnosis and microfluids.

There are several approaches designed to separate a population of homogeneous cells from the blood. These cell separation techniques can be grouped into two broad categories: (1) invasive procedures based on the selection of fixed and stained cells using various cell-specific markers; and (2) non-invasive procedures for the isolation of living cells using a specific biophysical parameter of a population of cells of interest.

Invasive techniques include fluorescence activated cell separation (FACS), magnetically activated cell separation (MACS) and immunomagnetic colloid separation. Normally, FACS is a positive selection technique that uses a fluorescently labeled marker to bind to cells that express a specific cell surface marker. FACS can also be used to permeabilize and stain cells for intracellular markers that can form the basis of separation. It is fast, normally works at a speed of 1,000 to 1,500 Hz and is well established in laboratory medicine. High rates of false positives are associated with the FACS due to the low number of photons obtained during extremely short residence times at high speeds. Complicated multiparameter separation approaches can be used to enhance the specificity of the FACS, but multi-analyte-based FACS may not be practical for routine clinical trials due to the high cost associated with it. The clinical application of FACS is also limited because it requires considerable operator expertise, is laborious, results in cell loss due to multiple manipulations and the cost of the equipment is prohibitive.

MACS is used as a cell separation technique in which cells expressing a specific surface marker are isolated from a sample of cells using magnetic beads coated with an antibody against the surface marker. The MACS has the advantage of being cheaper, easier and faster to perform compared to the FACS. It is affected by cell loss due to multiple manipulations and handling. In addition, magnetic beads are often autofluorescent and do not easily separate from cells. As a result, many of the immunofluorescence techniques used to investigate functions and cell structure are not compatible with this approach.

A magnetic colloid system has been used in the isolation of blood cells. This colloid system uses ferromagnetic nanoparticles that are coated with goat anti-mouse IgG that can easily bind to cell surface antigen specific monoclonal antibodies. Cells that are labeled with ferromagnetic nanoparticles align with a magnetic field along ferromagnetic Ni lines deposited by lithographic techniques on an optically transparent surface. This approach also requires multiple stages of cell handling that include mixing cells with magnetic beads and surface separation. It is also not possible to separate the individual cells from the sample for further analysis.

Non-invasive techniques include separation by charge flow, which employs a horizontal cross-flow liquid gradient opposite to an electric field in order to separate cells based on their characteristic surface charge densities. Although this approach can separate cells based purely on biophysical differences, it is not specific enough. There have been attempts to modify the characteristics of the device (eg, separator screens, buffer flow conditions, etc.) to address this major drawback of the technique. None of these device feature modifications has provided a practical solution given the expected individual variability in different samples.

Since prior art procedures are affected by high cost, low performance and lack of specificity, there is a need for a procedure to remove a particular cell type from a mixture that overcomes these limitations.

US 5,866,345 discloses a device for detecting the presence of an analyte in a liquid sample, the device comprising: a microfabricated solid substrate for defining: a sample inlet port; a mesoscale flow system comprising: a sample flow channel in fluid communication with said input port; and an analyte detection region in fluid communication with said flow channel comprising a binding moiety immobilized therein to specifically bind said analyte, said detection region having a mesoscale dimension; and a detection window arranged in said detection region to transmit a signal indicating the attachment of said analyte to a detection means disposed adjacent to said window.

US 2002/0115201 A1 discloses a microfluidic device comprising: a microwave monolithic integrated circuit (MMIC) within said microfluidic device for applying microwave radiation to a cavity defined by said microfluidic device.

US 6,613,525 B2 discloses an integrated microfluidic device having at least one microchannel formed in a generally flat substrate, said device comprising: at least one microchannel comprising an enrichment channel portion having at least one inlet and one outlet. ; and an enrichment medium present in said enrichment channel portion and containing specific binding moieties that bind to an objective of a sample such that at least a fraction of the objective is retained by said enrichment means; said enrichment channel portion is configured to (i) receive, through said input, the sample containing the objective and (ii) allow the movement of at least a part of the objective attached through said output.

US 6,344,326 B1 and US 6,074,827 respectively refer to a microfluidic process for purification and processing of nucleic acids. Microchannels or affinity capture receptors arranged in an enrichment zone can be used in the process.

Summary of the invention

According to one aspect, a microfluidic device is provided, which comprises:

(to)

 a first region of fixed obstacles arranged in a microfluidic channel defining a liquid flow path, in which the obstacles of the first region preferably bind to a first type of cell compared to a second type of cell, in which the Obstacles are placed in at least two columns and at least two rows, in which the rows are placed perpendicular to the liquid flow path in relation to the obstacles in the previous row, thereby forming a set of equilateral triangular obstacles; Y

(b)

 a second region of fixed obstacles arranged in the microfluidic channel, in which the obstacles of the second region preferably join a third type of cell compared to a fourth type of cell, in which the obstacles are placed in at least two columns and at least two rows, in which the rows are placed perpendicular to the liquid flow path and the obstacles of each successive row are displaced in a direction perpendicular to the liquid flow path in relation to the obstacles of the previous row , thus forming a set of equilateral triangular obstacles,

in which the second region is located beyond the first region in the microfluidic channel.

According to another aspect, a use of the device is provided as indicated in the independent claim. The dependent claims define embodiments.

The device can be used in procedures to separate cells from a sample (eg, to separate fetal red blood cells from maternal blood).

The procedure begins with the introduction of a sample that includes cells in one or more microfluidic channels. In one embodiment, the device includes at least two processing steps. For example, a mixture of cells is introduced into a microfluidic channel that selectively allows the passage of a desired cell type and the enriched cell population in the desired type is then introduced into a second microfluidic channel that allows the passage of the desired cell. to produce a more enriched cell population in the desired type. Cell selection is based on a property of the cells in the mixture, for example, size, shape, deformability, surface characteristics (e.g., cell surface receptors or antigens and membrane permeability) or intracellular properties (p eg expression of a specific enzyme).

In practice, the procedure can then be continued through a variety of processing steps using various devices. In one step, the sample is combined in the microfluidic channels with a solution that preferably smooths one type of cell compared to another type. At another stage, the cells are contacted with a device that contains obstacles in a microfluidic channel. Obstacles preferably link one type of cell compared to another type. Cells can also undergo separations based on size, deformability or shape. The methods of the invention can employ only one of the above steps or any combination of the steps, in any order, to separate the cells. Desirably, the methods of the invention recover at least 75%, 80%, 90%, 95%, 98% or 99% of the desired cells in the sample.

According to one embodiment, a microfluidic system is provided for the separation of a desired cell from a sample. This system may include devices for carrying out one or any combination of the steps of the procedures described above. One of these devices is a lysis device that includes at least two input channels; a reaction chamber (eg, a serpentine channel); and an output channel. The device may additionally include another inlet and a dilution chamber (eg, a serpentine channel). The lysis device is positioned so that at least two input channels are connected to the output through the reaction chamber. When a dilution chamber is present, it is disposed between the reaction chamber and the outlet and another inlet is disposed between the reaction chamber and the dilution chamber. The system may also include a cell impoverishment device that contains obstacles that preferably bind one type of cell compared to another type, e.g. eg, they are coated with anti-CD405, anti-CD35, anti-GPA or anti-CD71 antibodies. The system may also include a selection device that contains a two-dimensional set of locations to contain individual cells. The selection device may also contain actuators for the selective manipulation (eg, release) of individual cells in the assembly. Finally, the system may include a device for cell separation based on size. This device includes sieves that only allow the passage of cells below a desired size. The sieves are located with a microfluidic channel through which a cell suspension passes, as described herein. When used in combination, the system devices may be in liquid communication with each other. Alternatively, samples that pass through one device can be collected and transferred to another device.

By an "impoverished cell population" is meant a population of cells that has been processed to reduce the relative population of a specified cell type in a mixture of cells. Consequently, collecting those cells removed from the mixture also leads to a sample enriched in the cells removed.

By a "rich cell population" is meant a population of cells that has been processed to increase the relative population of a specified cell type in a mixture of cells.

By "lysis buffer" is meant a buffer that, when contact is made with a population of cells, will cause lysis of at least one type of cell.

By "causing lysis" it is meant that at least 90% of cells of a specific type are lysed.

By "non-lysate" it is meant that less than 10% of cells of a particular type are lysed. Desirably, less than 5%, 2% or 1% of these cells are lysed.

By "type" of cell is meant a population of cells that have a common property, e.g. eg, the presence of a specific surface antigen. A single cell can belong to several different cell types.

By "serpentine channel" is meant a channel that has a total length that is larger than the linear distance between the points of the ends of the channel. A serpentine channel can be oriented completely vertically or horizontally. Alternatively, a serpentine channel may be in "3D," p. eg, parts of the channel are oriented vertically and parts are oriented horizontally.

By "microfluidic" is meant that it has one or more dimensions of less than 1 mm.

By "rest of union" is meant a chemical species to which a cell binds. A binding moiety can be a compound coupled to a surface or the material that constitutes the surface. Exemplary binding moieties include antibodies, oligo-or polypeptides, nucleic acids, other proteins, synthetic polymers and carbohydrates.

By "obstacle" is meant an impediment to the flow in a channel, e.g. eg, a protrusion of a surface.

By "specifically linking" a type of cell is meant to unite cells of that type by a specified mechanism, e.g. eg, antibody-antigen interaction. The strength of the junction is generally sufficient to prevent detachment by the flow of liquid present when the cells are attached, although occasionally individual cells can detach under normal operating conditions.

By "rows of obstacles" is meant a series of obstacles placed so that the centers of the obstacles are placed in a substantially linear manner. The distance between rows is the distance between the lines of two adjacent rows in which the centers are located.

By "obstacle columns" is meant a series of obstacles placed perpendicular to a row so that the centers of the obstacles are placed substantially linearly. The distance between columns is the distance between the lines of two adjacent columns in which the centers are located.

The procedures are capable of separating specific cell populations from a complex mixture without fixing and / or staining them. As a result of obtaining live homogenous cell populations, many functional tests can be performed on the cells. The microfluidic devices described herein provide a simple, selective approach to processing cells. Other features and advantages of the invention will be apparent from the following description and claims.

Brief description of the drawings

Figure 1 is a schematic design of a microfluidic device that allows selective lysis of cells.

Figure 2 is an illustration of the channel design for the introduction of three liquids into the device, e.g. eg, blood sample, lysis buffer and diluent.

Figure 3 is an illustration of a repeat unit of the reaction chamber of the device where a sample of cells is passively mixed with a lysis buffer. In one example, 133 units are connected to form the reaction chamber.

Figure 4 is an illustration of the output channels of the device.

Figure 5 is an illustration of a device for cell lysis.

Figures 6A and 6B are illustrations of a process for manufacturing a device of the invention.

Figure 7 is a schematic diagram of a cell binding device.

Figure 8 is an exploded view of a cell binding device.

Figure 9 is an illustration of the obstacles of a cell binding device.

Figure 10 is an illustration of types of obstacles.

Figure 11A is a schematic representation of a square set of obstacles. The square set has a capture efficiency of 40%. Figure 11B is a schematic representation of an equilateral triangular set of obstacles. The equilateral triangular set has a capture efficiency of 56%.

Figure 12A is a schematic representation of the calculation of hydrodynamic efficiency for a square set. Figure 12B is a schematic representation of the calculation of hydrodynamic efficiency for a diagonal matrix.

Figures 13A-13B are graphs of hydrodynamic (13A) and overall (13B) efficiency for a square and triangular assembly for a pressure drop of 150 Pa / m. This pressure drop corresponds to a flow rate of 0.75 ml / h in the flat geometry.

Figure 14A is a graph of overall efficiency as a function of pressure drop. Figure 14B is a graph of the effect of separation of obstacles on average speed.

Figure 15 is a schematic representation of the placement of obstacles for a higher capture efficiency for an equilateral triangular set of obstacles in a stepped set. The capture radius is rcap2 = 0.339 /. The obstacles are numbered so that the first number refers to the triangle number and the second number refers to the vertex of the triangle. The stepped set has a capture efficiency of 98%.

Figure 16A is a graph of the percentage of cell capture as a function of the flow rate for an obstacle geometry of 100 µm in diameter with a spacing of 50 µm between its edges. The operating flow regime was established through several cell types: cancer cells, normal connective tissue cells and maternal and fetal samples. An optimal workflow regime is 2.5 ml / h. Figure 16B is a graph of the percentage of cell capture as a function of the ratio of target cells and white blood cells. The model system was generated by adding a defined number of cancer cells, normal connective tissue cells or umbilical cord blood cells in a defined number of leukocyte layer cells of adult blood. The ratio of contaminating cells and target cells gradually increased in 5 log with only 10 target cells in the mixture. The yield was calculated as the difference between the number of added cells captured in positions and the number of cells added to the sample.

Figure 17 is an illustration of various views of the input and outputs of a cell binding device.

Figure 18 is an illustration of a manufacturing method of a cell binding device.

Figure 19 is an illustration of a mixture of cells flowing through a cell binding device.

Figure 20A is an illustration of a cell binding device for trapping different types of cells in series. Figure 20B is an illustration of a cell binding device for trapping different cell types in parallel.

Figure 21 is an illustration of a cell binding device that allows recovery of bound cells.

Figure 22A is an optical micrograph of fetal red blood cells adhered to an obstacle of the invention. Figure 22B is a fluorescent micrograph showing the results of an FISH analysis of a fetal red blood cell attached to an obstacle of the invention. Figure 22C is an enlarged micrograph of Figure 22B showing the individual hybridization results for the fetal red blood cell.

Figure 23 is an illustration of a cell binding device in which beads trapped in a hydrogel are used to capture cells.

Figure 24A is an illustration of a device for size-based separation.

Figure 24B is an electron micrograph of a device for size-based separation.

Figure 25 is a schematic representation of a device of the invention for isolating and analyzing fetal red blood cells.

The figures are not necessarily to scale.

5 Detailed description of the invention

The invention features devices for use in methods for separating a desired cell from a mixture or enriching the population of a desired cell in a mixture. The procedures are generally based on sequential processing steps, each of which reduces the number of unwanted cells in the mixture, but a processing step can be used in the procedures. The devices for carrying out 10 different processing steps can be separated or integrated into a microfluidic system. The devices of the invention include a device for cell binding. In one embodiment, processing steps are used to reduce the number of cells before selecting them. Desirably, the procedures retain at least 75%, 80%, 90%, 95%, 98% or 99% of the cells compared to the initial mixture, while potentially enriching the population of desired cells by a factor of at least 100, 1000,

15 10,000, 100,000 or even 1,000,000 in relation to one or more unwanted cell types. The procedures can be used to separate or enrich cells that circulate in the blood (Table 1).

Table 1: Types, concentrations and sizes of blood cells.

Cell type

Concentration (cells / µl) Size (µm)

Red blood cells (GR)

4.2-6.1 x 106 4-6

Segmented Neutrophils (GB)

3600 > 10

Band Neutrophils (GB)

120 > 10

Lymphocytes (GB)

1500 > 10

Monocytes (GB)

480 > 10

Eosinophils (GB)

180 > 10

Basophils (GB)

120 > 10

Platelets

500 x 10-3 1-2

Fetal nucleated red blood cells

2 -50 x 10-3 8-12

Dispositives

20 A. Cellular Lysis

A device is used to lyse a population of cells selectively, e.g. eg, maternal red blood cells, in a mixture of cells, e.g. eg, maternal blood. This device allows the processing of large amounts of cells under almost identical conditions. Desirably, the lysis device removes a large number of cells before further processing. Waste, p. eg, cell membranes and proteins, can be trapped, e.g. eg

25 by filtration or precipitation, before any further processing.

Device. A design for a lysis device is shown in Figure 1. The overall branched architecture of the channels of the device allows equivalent pressure drops along the parallel processing networks. The device can be functionally separated into four distinct sections: 1) distributed input channels that carry liquids, e.g. eg, blood, lysis reagent and wash buffer, at junctions 1 and 2 (Figure 2); 2) a camera

30 serpentine reaction for the cell lysis reaction that is housed between the two junctions (Figure 3); 3) a dilution chamber beyond Union 2 for dilution of the lysis reagent (Figure 3); and 4) distributed outflow channels that carry the lysed sample to a collection vial or other microfluidic device (Figure 4).

Input / output channels. Branched channel input and output networks allow for uniform distribution of reagents across all channels (8, as shown in Figure 1). The three ports for connecting the macro world with the device usually have a diameter that varies between 1 mm -10 mm, p. e.g., 2, 5, 6 or 8 mm. Seals can be formed with ports 1, 2 and 3, p. e.g., through an external collector integrated with the device (Figure 1). The three solution vials, p. eg, blood, lysis reagent and diluent, can be connected to such a manifold. The input channels of ports 1, 2 and 3 to the reaction and mixing chambers, for the three solutions shown in Figure 1, can be separated in the z-plane of the device (three layers, each with a set of distribution channels, see Figure 2) or stay in the external manifold. If they stay

In the external manifold, the distribution channels are, for example, CNC (numerically controlled by computer), machined in stainless steel and can have dimensions of 500 µm in diameter. The manifold can be tightly connected to the device in ports that are recorded in the 1 ', 2' and 3 'locations shown in Figure 1. Locating the distribution channels in a manifold reduces the complexity and cost of the device. Keeping the distribution channels in the device will allow greater flexibility to select a smaller channel size, while avoiding any problems of contamination by transfer between samples. Each sample input channel can have a separate output or, as shown in Figure 4, the output channels for each sample input are combined. As an alternative to a collector, tubes for each liquid inlet or outlet can be attached to the device, e.g. eg, by compression fitting to joints or nozzles or using water tight connections such as a luer seal. The channels of the device that carry the liquids to the joints and subsequent mixing chambers can have a width and depth ranging from 10 µm -500 µm, p. eg, a maximum width and depth of 10 µm, 25 µm, 50 µm, 75 µm, 100 µm, 150 µm, 200 µm, 250 µm, 350 µm or 450 µm. Desirably, the architecture of the channel is rectangular but it can also be circular, semicircular, V-shaped or any other appropriate shape. In one implementation, the output channel (or channels) has a cross-sectional area equal to the sum of the cross-sectional areas of the input channels.

Reaction and dilution chambers. For lysis and dilution, two liquid streams are combined and allowed to pass through the chambers. The cameras can be serpentine or linear channels. In the device depicted in Figure 1, the sample and lysis buffer are combined at junction 1 and the lysed sample and diluent are combined at junction 2. The serpentine architecture of the reaction chamber and dilution chamber allows sufficient residence time of the two reactive solutions for proper mixing by diffusion or other passive mechanisms, while maintaining a reasonable overall size for the device (Figure 3). The serpentine channels can be constructed in 2D or 3D, e.g. eg, to reduce the total length of the device or to introduce chaotic advection for enhanced mixing. For short residence times, a linear camera may be desired. Exemplary residence times include at least 1 second, 5 seconds, 10 seconds, 30 seconds, 60 seconds, 90 seconds, 2 minutes, 5 minutes, 30 minutes, 1 hour or greater than 1 hour. The flow rate of the reaction / dilution chambers can be precisely controlled by controlling the effective width, depth and length of the channels to allow sufficient mixing of the two reagents while at the same time allowing optimal processing performance. In one implementation, the serpentine mixing chambers for cell lysis (reaction chamber) and for dilution of the lysed sample (dilution chamber) have a liquid volume of 26 µl each. Other examples of reaction / dilution chamber volumes range from 10-200 µl, e.g. eg, maximum 20, 50, 100 or 150 µl. In some implementations, the width and depth of the reaction and dilution chambers have the same range as the input and output channels, that is, 10 to 500 µm. Alternatively, the chambers may have a cross-sectional area equal to the combined areas of all input (or output) channels in order to ensure a uniform flow rate through the device. In one example, the cameras are 100 µm x 100 µm channels. The total length of the cameras can be at least 1 cm, 5 cm, 10 cm, 20 cm, 30 cm, 40 cm or 50 cm.

For maternal GR lysis, the flow rates of the device can vary from processing 5 -16 µl of blood per second resulting in a processing time of 20 -60 minutes for 20 ml samples or a processing time of 10 -30 min for 10 ml samples. The sample volume required to capture a sufficient number of fetal cells is expected to be less than 10 ml due to the effectiveness of the procedure. Thus, the performance of the device per sample is expected to be less than 10 minutes. A residence time of> 30 seconds from the moment of convergence of the two solutions, maternal blood and lysis reagent, within the passive mixer is considered sufficient to obtain an effective hemolysis (T. Maren, Mol. Pharmacol. 1970, 6 : 430). Alternatively, the concentration of lysis reagent can be adjusted to compensate for the residence time in the reaction chamber. The flow rates and residence times for other cell types can be determined theoretically or by experimentation. In one implementation, the flows in each channel are limited to <20 µl / s to ensure that the shear stress of the walls on the cells is less than 1 dyne / cm2 (it is known that a shear stress> 1 dyne / cm2 it affects the cells functionally, although no harmful effects are observed in most cells until it exceeds 10 dynes / cm2). In one implementation, the flow rate in each channel is a maximum of 1, 2, 5, 10, 15 µl / s. Referring to Figure 1, the effective length of the diluent inlet channel leading to the junction 2 may be shorter than the effective length of the reaction chamber. This feature allows the diluent to flow to and feed the channels after junction 2, before the arrival of the lysed sample at junction 2. The excess buffer previously collected in the outlet vial can act as a secondary diluent of the sample lysed when collected, for example, for further processing or analysis. Additionally, the diluent feeds the post-junction 2 channels to allow smoother flow and fusion of the lysed sample with the buffer in the dilution chamber, and this feed eliminates any effects of harmful surface tension of dry channels in the sample lysate The diameter of the channels carrying the diluent can be adjusted to allow the diluent to reach junction 2 at the same time as the lysed blood to avoid any problems associated with forced air from the reaction chamber as the sample and buffers are introduced of lysis.

Although the above description focuses on a device with eight parallel processing channels, any number of channels can be used, e.g. e.g., 1, 2, 4, 16 or 32, depending on the size of the device. The device is described in terms of combining two liquids for lysis and dilution, but three or more liquids for lysis or dilution can be combined. The combination can be in a joint or a series of joints, e.g. eg, to control the rate of sequential reagent addition. Additional liquid inlets may be added, e.g. eg, to functionalize the remaining cells, alter the pH or cause precipitation of undesirable components. In addition, the exact geometry and dimensions of the channels can be altered (exemplary dimensions are shown in Figure 5). The devices of the invention can be disposable or reusable. Disposable devices reduce the risk of contamination between samples. Reusable devices may be desirable in some cases and the device may be cleaned, e.g. eg, with various detergents and enzymes, e.g. eg, proteases or nucleases, to avoid contamination.

Pumping. In one implementation, the device employs negative pressure pumping, e.g. eg, using syringe pumps, peristaltic pumps, vacuum cleaners or vacuum pumps. The negative pressure allows the processing of the total volume of a clinical blood sample, without leaving an unprocessed sample in the channels. Positive pressure can also be used, e.g. eg, from a syringe pump, peristaltic pump, displacement pump, liquid column or other liquid pump, to pump samples through a device. Sample loss due to dead volume problems related to positive pressure pumping can be overcome by expelling the residual sample with buffer. The pumps are normally connected to the device through airtight seals, e.g. eg, using silicone gaskets.

The liquid flow rates of parallel channels of the device can be controlled together or separately. Variable and differential control of the flow rates of each channel can be achieved, for example, by using an individually controllable multi-channel syringe collector. In this implementation, the distribution of the input channel will be modified so that it is decoupled from all parallel networks. The outlet can collect the output of all channels through a single manifold connected to a suction (a tight seal is not required) that has an outlet to a collection vial or another microfluidic device. Alternatively, the output of each network can be collected separately for further processing. Separate inputs and outputs allow parallel processing of multiple samples from one or more individuals.

Manufacturing. A variety of techniques can be employed to manufacture a device and the technique employed will be selected based in part on the chosen material. Exemplary materials for manufacturing the devices of the invention include glass, silicon, steel, nickel, poly (methyl methacrylate) (PMMA), polycarbonate, polystyrene, polyethylene, polyolefins, silicones (e.g., poly (dimethylsiloxane)) and combinations thereof. . Other materials are known in the art. Procedures for manufacturing channels of these materials are known in the art. These procedures include, photolithography (e.g., stereolithography or x-ray photolithography), molding, embossing, silicon micromachining, wet or dry chemical etching, grinding, diamond cutting, Lithographie Galvanoformung and Abformung (LIGA) and electroplating. For example, for glass, traditional silicon photolithography manufacturing techniques followed by wet (KOH) or dry etching (fluoride reactive ion or other reactive gas) can be used. Techniques such as laser micromachining for plastic materials with high photon absorption efficiency can be adopted. This technique is suitable for lower performance manufacturing due to the serial nature of the procedure. For plastic devices produced in series, thermoplastic injection molding and compression molding are suitable. Conventional thermoplastic injection molding used for mass production of compact discs (which maintains the fidelity of the characteristics in submicras) can also be used to manufacture the devices. For example, the characteristics of the device are duplicated in a glass pattern by conventional lithophotography. The glass pattern is electroformed to provide a resistant mold, resistant to thermal shock, thermoconductive and hard. This mold serves as the template template for injection molding or compression molding of the features in a plastic device. Depending on the plastic material used to manufacture the devices and the requirements of optical quality and performance of the finished product, compression molding or injection molding can be chosen as the manufacturing process. Compression molding (also called hot embossing or embossing) has the advantages of being compatible with high molecular weight polymers, which are excellent for small structures, but it is difficult to use to duplicate structures with high aspect ratio and has longer duration cycles. Injection molding works well for high aspect ratio structures, but is more suitable for low molecular weight polymers.

A device can be manufactured in one or more pieces that are mounted later. In one embodiment, the separate layers of the device contain channels for a single fluid, as in Figure 1. The layers of a device can be joined together with tweezers, adhesives, heat, anodic bonding or reactions between surface groups (e.g. ., union of wafers). Alternatively, a device with channels in more than one -plane can be manufactured as a single piece, e.g. eg, using stereolithography or other three-dimensional manufacturing techniques.

In one implementation, the device is made of PMMA. The characteristics, for example, those shown in Figure 1, are transferred to an electroformed mold using standard photolithography followed by electroplating. The mold is used for hot embossing the characteristics in the PMMA at a temperature close to its glass transition temperature (105 ºC) under pressure (5 to 20 tons) (the pressure and temperature will be adjusted to allow high doubling fidelity of the deepest feature of the device) as shown in Figure 6A. Then, the mold is cooled to allow removal of the PMMA device. A second piece used to close the device, composed of similar or different material, can be attached to the first piece using vacuum assisted thermal bonding. The vacuum prevents the formation of air voids in the junction regions. Figure 6B shows a cross section of the two-piece device assembly at the junction of Port 1 (source for blood sample) and the feeding channel.

Chemical derivatization. To reduce the non-specific adsorption of cells or compounds released by lysed cells on the canal walls, one or more canal walls can be chemically modified to be non-adherent or repulsive. The walls can be coated with a thin film coating (eg, a monolayer) of commercial non-stick reagents, such as those used to form hydrogels. Additional examples of chemical species that can be used to modify canal walls include oligoethylene glycols, fluorinated polymers, organosilanes, thiols, polyethylene glycol, hyaluronic acid, bovine serum albumin, polyvinyl alcohol, mucin, poly-HEMA, PEG-methacrylate and agarose. Loaded polymers can also be used to repel species of opposite charge. The type of chemical species used for repulsion and the procedure for joining the canal walls will depend on the nature of the species that repel and the nature of the walls and the species that bind. Such surface modification techniques are well known in the art. The walls can be functionalized before or after mounting the device.

Channel walls can also be coated in order to capture sample materials, e.g. eg, membrane fragments or proteins.

Procedures In the present implementation, a sample of cells is introduced, e.g. eg, maternal blood, in one or more microfluidic channels. A lysis buffer containing reagents for selective lysis of a population of sample cells is then mixed with the blood sample. Desirably, mixing occurs by passive means, e.g. eg, diffusion or chaotic advection, but active means can be employed. Additional passive and active mixers are known in the art. The lysis reaction is allowed to continue for a desired period of time. This period of time can be controlled, for example, by the length of the channels or by the flow of the liquids. In addition, it is possible to control the volumes of the mixed solutions in the channels by altering the relative volumetric flow rates of the solutions, e.g. eg, altering the channel size or flow rate. The flow can be slowed, increased or stopped for any desired period of time. After lysis has occurred, a diluent may be introduced into the channel in order to reduce the concentration of lysis reagents and any potentially harmful species (eg, endosomal enzymes) released by lysed cells. The diluent may contain species that neutralize lysis reagents that otherwise alter the liquid's environment, e.g. eg, pH or viscosity, or it may contain reagents for surface or intracellular labeling of cells. The diluent can also reduce the optical density of the solution, which may be important for some detection schemes, e.g. eg absorbance measures.

Exemplary cell types that can be lysed using the procedures described herein include adult red blood cells, white blood cells (such as T lymphocytes, B lymphocytes and helper T lymphocytes), infected white blood cells, tumor cells and infectious organisms (eg. , bacteria, protozoa and fungi). The lysis buffer for these cells can include cell-specific IgM molecules and complement cascade proteins to initiate complement-mediated lysis. Another class of lysis buffer may include viruses that infect a specific cell type and cause lysis as a result of duplication (see, eg, Pawlik et al. Cancer 2002, 95: 1171-81). Other lysis buffers are known in the art.

A device for the selective lysis of maternal red blood cells (GR) can be used to enrich a blood sample in fetal cells. In this example, a sample of maternal blood, 10-20 ml, is processed within the first one to three hours after sample collection. If the processing is delayed beyond three hours, the sample can be stored at 4 ° C until it is processed. The lysis device allows mixing the lysis reagent (NH4Cl (0 to 150 mM) + NaHCO3 (0.001 to 0.3 mM) + acetazolamide (0.1 to 100 µM)) with the maternal blood to allow selective lysis of the maternal red blood cells for the underlying principle of the Orskov-Jacobs-Stewart reaction (see, for example, Boyer et al. Blood 1976, 47: 883-897). The high selective permeability of the carbonic anhydrase inhibitor, acetazolamide, in fetal cells allows selective hemolysis of the maternal red blood cells. The endogenous carbonic anhydrase of the maternal cells converts HCO3-into carbon dioxide, which smooths the maternal red blood cells. The enzyme is inhibited in fetal red blood cells and those cells do not lyse. A diluent (e.g., phosphate buffered saline) may be added after a period of contact between the lysis reagents and the cell sample to reduce the risk of a portion of the fetal red blood cells licking after exposure prolonged to reagents.

B. Cell junction

Another device of the invention involves the removal of whole cells from a mixture by joining the cells to the surfaces of the device. The surfaces of such a device contain substances, e.g. eg, antibodies or ligands for cell surface receptors, which bind a specific subpopulation of cells. This step of the procedure can employ positive selection, that is, the desired cells are attached to the device, or it can employ negative selection, that is, the desired cells pass through the device. In both cases, the population of cells containing the desired cells is collected for analysis or further processing.

Device. The device is a microfluidic flow system that contains a set of obstacles in various ways that are capable of joining a population of cells, e.g. eg, those that express a specific surface molecule of a mixture. The bound cells can be analyzed directly in the device or removed from the device, e.g. eg, for analysis or further processing. Alternatively, cells not bound to obstacles can be collected, e.g. eg, for analysis or further processing.

An exemplary device is a flow apparatus that has a flat plate channel through which cells flow; Such a device is described in US Pat. UU. No. 5,837,115. Figure 7 shows an exemplary system that includes an infusion pump to perfuse a mixture of cells, e.g. eg, blood, through the microfluidic device. Other pumping procedures may be employed, as described herein. The device may be optically transparent, or have transparent windows, to visualize the cells during the flow through the device. The device contains obstacles distributed, in an orderly set, along the flow chamber. Desirably, the upper and lower surfaces of the device are parallel to each other. This concept is represented in Figure 8. Obstacles can be part of the lower or upper surface and, desirably, define the height of the flow channel. It is also possible that a part of the obstacles are located on the lower surface and the rest on the upper surface. Obstacles may be in contact with both upper and lower surfaces of the chamber or there may be a gap between an obstacle and a surface. Obstacles may be coated with a binding residue, e.g. eg, an antibody, a charged polymer, a molecule that binds a cell surface receptor, an oligo-or polypeptide, a viral or bacterial protein, a nucleic acid or a carbohydrate, that binds to a population of cells, p. eg, those expressing a specific surface molecule of a mixture. Other binding moieties that are specific to a particular cell type are known in the art. In an alternative embodiment, the obstacles are made of a material to which a specific type of cell is attached. Examples of such materials include organic polymers (charged or unloaded) and carbohydrates. Once a binding residue is coupled to the obstacles, a coating, as described herein, can also be applied to any exposed surface of the obstacles to avoid

20 non-specific adhesion of cells to obstacles.

An obstacle geometry is shown in Figure 9. In one example, the obstacles are engraved on a surface area of 2 cm x 7 cm on a substrate with overall dimensions of 2.5 cm x 7.5 cm. A margin of 2 mm is left around the substrate to attach it to the upper surface to create a closed chamber. In one embodiment, the diameter of the obstacles is 50 µm with a height of 100 µm. The obstacles may be placed in a two-dimensional rowset with a distance of 100 µm between the centers. This placement provides 50 µm openings for cells to flow between obstacles without being mechanically compressed or damaged. Desirably, the obstacles in a row are displaced, e.g. eg, 50 µm with respect to adjacent rows. This alternate pattern can be repeated throughout the design to ensure an increased collision frequency between cells and obstacles. The diameter, width or length of the obstacles may be at

30 minus 5, 10, 25, 50, 75, 100 or 250 µm and a maximum of 500, 250, 100, 75, 50, 25 or 10 µm. The spacing between the obstacles can be at least 10, 25, 50, 75, 100, 250, 500 or 750 mm and a maximum of 1000, 750, 500, 250, 100, 75, 50 or 25 µm. Table 2 lists exemplary spacing based on the diameter of the obstacles.

Table 2. Exemplary spacing for obstacles.

Obstacle diameter (µm)

Distance between obstacles (µm)

100

fifty

100

25

fifty

fifty

fifty

25

10

25

10

fifty

10

fifteen

35 The dimensions and geometry of obstacles can vary significantly. For example, obstacles may have cylindrical or square cross sections (Figure 10). The distance between obstacles may also vary and may be different in the direction of the flow compared to the direction orthogonal to the flow. In some embodiments, the distance between the edges of the obstacles is slightly larger than the size of the largest cell in the mixture. This placement allows the flow of cells without being compressed

40 mechanically between the obstacles and, therefore, are damaged during the flow process, and also maximizes the number of collisions between cells and the obstacles in order to increase the probability of binding. The flow direction with respect to the orientation of the obstacles can also be altered to enhance the interaction of cells with obstacles.

Exemplary obstacle placements are shown in Figures 11A-11B. All these placements have a calculated capture efficiency. The calculation of the cell junction considered two different geometries: a square set (Figure 11A) and an equilateral triangular set (Figure 11B). In general, the results are presented in terms of the effectiveness of adhesion. The calculations consist of two parts, calculate the hydrodynamic efficiency (

f) and the probability of adhesion. Hydrodynamic efficiency was determined as the ratio of the capture radius and half the distance between the cylinders (Figures 12A and 12B). For the square set, f = (2rcap / l) * 100%, and for other sets, f = ((rcap1 + rcap2) / d1) * 100%, where d1 = d2 = l / N2 for a diagonal square set, and d1 = lN3 / 2, d2 = l / 2 for a triangular set. The probability of adhesion represents the fraction of cells that can withstand the force applied to the cell assuming an average of 1.5 junctions per cell and 75 pN per junction.

For the triangular set, more cells adhered to the second set of obstacles than to the first set. Figures 13A-13B show that efficiency is reduced as the spacing between obstacles increases. As spacing increases there is a larger region outside the capture radius and the cells never come in contact with obstacles. In addition, for the analyzed flows (0.25 -1 ml / h), the overall probability of adhesion is high because the force per cell is less than the force to break the joints. For a triangular set and 150 micrometer spacing, the overall capture efficiency drops by 12% when the flow rate increases from 0.25 to 1 ml / h (Figures 14A-14B). Adhesion does not improve by going to lower flow rates, since hydrodynamic capture does not improve. The average speed increases as the spacing between obstacles increases. The reason for this is that the calculations used a constant pressure drop. This differs from experiments in which the flow rate remains fixed and the pressure drop varies. The results can be extrapolated from one case to another by one skilled in the art.

A repetitive triangular set provides limited capture of target cells because most of the capture occurs in a few first rows. The reason for this is that the flow field is established in these rows and repeated. The first capture radius does not produce much capture while the majority of the capture is within the second capture radius (Figure 15). Once the cells within the capture radii are captured, the only way in which the capture could occur is through cell-cell collisions to move cells out of their current lines or secondary capture. Referring to Figure 15, in order to enhance the capture, after the flow field is established, the rows move a distance in the vertical direction (perpendicular to the flow) by a distance equal to rrap2 = 0.339l. The first five columns form two regular regions of equilateral triangles. This allows the flow to be established and compatible with the solution for an equilateral triangular assembly. To promote the capture of cells that fall outside of rcap2, the fourth column is displaced a distance rcap2. All columns are separated by a distance equal to l / 2. It is shown that a cell that falls outside of rcap2 is captured by the first obstacle of the fourth triangle. Triangles 4 and 5 would be equilateral. In triangle 6, vertex 3 is shifted down a distance rcap2. This placement can be repeated every three triangle, that is, the repetition distance is 2.5l. Figures 16A and 16B illustrate the capture efficiency as a function of the flow rate and relative population of the desired cells.

Desirably, the top layer is made of glass and has two ultrasonic perforated grooves for the inflow and outflow. The inlet / outlet dimensions of the slit are, for example, 2 cm long and 0.5 mm wide. Figure 17 shows the details of the input / output geometry. A collector can then be incorporated over the inlet / outlet grooves. The inlet manifold accepts blood cells from an infusion syringe pump or any other delivery vehicle, for example, through a flexible, biocompatible tube. Similarly, the outlet manifold is connected to a reservoir to collect the solution and the cells that leave the device.

The configuration and geometry of input and output can be designed in various ways. For example, circular inputs and outputs can be used. An entry region without obstacles is then incorporated into the design to ensure that blood cells are distributed evenly when they reach the region where the obstacles are located. Similarly, the outlet is designed with an unobstructed exit region to collect cells that leave evenly without damaging them.

The overall size of an exemplary device is shown in Figure 9 (upper diagram). The length is 10 cm and the width is 5 cm. The area that is covered with obstacles is 9 cm x 4.5 cm. The design is flexible enough to accommodate larger or smaller sizes for different applications.

The overall size of the device may be smaller or larger, depending on the flow performance and the number of cells to be removed (or captured). A larger device could include a greater number of obstacles and a larger surface area for cell capture. Such a device may be necessary if the quantity of sample, e.g. eg, blood, which is going to be processed is large.

Manufacturing. An exemplary method for manufacturing a device of the invention is summarized in Figure 18. In this example, standard photolithography is used to create a pattern of photoresist obstacles on a silicon wafer on insulator (SOI). An SOI wafer consists of a 100 µm thick Si (100) layer on top of a 1 µm thick SiO2 layer on a 500 µm thick Si (100) wafer. To optimize the adhesion of the photoresist, the SOI wafers can be exposed to high temperature hexamethyldisilazane vapors before the photoresist coating. The wafer is coated by centrifugation with the UV-sensitive photoresist, baked for 30 minutes at 90 ° C, exposed to UV light for 300 seconds through a chrome contact mask, revealed for 5 minutes in a developer and post-bake for 30 minutes at 90 ° C. The procedure parameters can be altered depending on the nature and thickness of the photoresist. The pattern of the chrome contact mask is transferred to the photoresist and determines the geometry of the obstacles.

After the formation of the photoresist pattern that is the same as the obstacle pattern, engraving begins. SiO2 can serve as a stop for the engraving procedure. You can also control the engraving stop at a certain depth without the use of a stopper layer. The photoresist pattern is transferred to the Si layer 100 µm thick in a plasma recorder. Deep multiplex engraving can be used to achieve uniform obstacles. For example, the substrate is exposed for 15 seconds to an SF6 that flows in fluorine-rich plasma and then the system is changed to a C4F8 only that flows in fluorocarbon-rich plasma for 10 seconds, which covers all surfaces with a protective film . In the subsequent etching cycle, exposure to ion bombardment preferably removes the polymer from horizontal surfaces and the cycle is repeated several times until, e.g. eg, the SiO2 layer is reached.

To couple a binding residue to the surfaces of the obstacles, the substrate can be exposed to an oxygen plasma before surface modification to create a layer of silicon dioxide, to which the binding remains can be attached. Then, the substrate can be rinsed twice with deionized, distilled water and allowed to air dry. Silane immobilization on exposed glass is performed by immersing samples for 30 seconds in freshly prepared 2% v / v solution of 3 - [(2-aminoethyl) amino] propyltrimethoxysilane in water followed by further washing in deionized, distilled water. Then, the substrate is dried in nitrogen gas and baked. Then, the substrate is immersed in 2.5% v / v solution of glutaraldehyde in phosphate buffered saline for 1 hour at room temperature. The substrate is then rinsed again and immersed in a 0.5 mg / ml solution of binding residue, e.g. eg, anti-CD71, anti-CD36, anti-GPA or anti-CD45, in deionized water, distilled for 15 minutes at room temperature to couple the binding agent to the obstacles. The substrate is then rinsed twice in deionized, distilled water and soaked overnight in 70% ethanol for sterilization.

There are several different techniques from those of the procedure described above by means of which binding remains can be immobilized on the obstacles and the surfaces of the device. A simple physioabsorption on the surface can be the choice for its simplicity and cost. Another approach may use monolayers grouped together (eg, thiols on gold) that are functionalized with various binding moieties. Additional procedures may be used depending on the junction moieties being joined and the material used to make the device. Surface modification methods are known in the art. In addition, some cells may preferentially bind to the unaltered surface of a material. For example, some cells may preferentially bind to positively charged, negatively charged surfaces or hydrophobic surfaces or to chemical groups present in some polymers.

The next stage involves the creation of a flow device by joining a top layer to the microfabricated silicon that contains the obstacles. The upper substrate may be glass to provide visual observation of cells during and after capture. Thermal bonding or a UV curable epoxy can be used to create the flow chamber. The upper and lower parts can also be adjusted by compression using, for example, a silicone gasket. Such compression adjustment can be reversible. Other binding methods (eg, wafer bonding) are known in the art. The procedure used may depend on the nature of the materials used.

The cell binding device can be made of different materials. Depending on the choice of material, different manufacturing techniques can also be used. The device may be made of plastic, such as polystyrene, using a hot embossing technique. Obstacles and other necessary structures are embossed in the plastic to create the bottom surface. Then an upper layer can be attached to the lower layer. Injection molding is another approach that can be used to create such a device. Soft lithography can also be used to create a camera made entirely of poly (dimethylsiloxane) (PDMS), or only obstacles in PDMS can be created and then attached to a glass substrate to create the closed chamber. Another approach involves the use of epoxy casting techniques to create obstacles through the use of UV or temperature curable epoxy on a pattern that has the negative replica of the intended structure. The laser or other types of micromachining approaches can also be used to create the flow chamber. Other suitable polymers that can be used in the manufacture of the device are polycarbonate, polyethylene and poly (methyl methacrylate). In addition, metals such as steel and nickel can also be used to make the device of the invention, e.g. eg, by traditional metal machining. Three-dimensional manufacturing techniques (eg, stereolithography) can be used to make a device in one piece. Other manufacturing processes are known in the art.

Procedures Methods in which a device of the invention can be used involve contacting a mixture of cells with the surfaces of a microfluidic device. A population of cells of a complex mixture of cells such as blood is then attached to the surfaces of the device. Desirably, at least 60%, 70%, 80%, 90%, 95%, 98% or 99% of the cells that are capable of binding to the surfaces of the device are removed from the mixture. The surface coating is desirably designed to minimize non-specific cell binding. For example, at least 99%, 98%, 95%, 90%, 80% or 70% of cells unable to bind to the rest of the junction do not bind to the surfaces of the device. Selective binding in the device results in the separation of a specific living cell population from a mixture of cells. Obstacles are present in the device to increase the surface area so that the cells interact with them while they are in the chamber containing the obstacles, so that the probability of binding increases. The flow conditions are such that the cells are handled very smoothly in the device without the need to deform mechanically in order to enter between the obstacles. Positive pressure pumping can be used

or negative pressure or flow from a liquid column to transport cells into and out of the microfluidic devices of the invention. In an alternative embodiment, the cells are separated from a non-cellular matter, such as a non-biological matter (e.g., beads), non-viable cell debris (e.g., membrane fragments) or molecules (e.g. ., proteins, nucleic acids or cell lysates).

Figure 19 shows cells expressing a surface antigen that binds to a binding residue coated on obstacles, while other cells flow through the device (the small arrow on the cells represents the directionality of the cells that are not bound to the surface). The upper and lower surfaces of the flow apparatus may also be coated with the same binding residue, or a different binding residue, to promote cell binding.

Exemplary cell types that can be separated using the procedures described herein include adult red blood cells, fetal red blood cells, trophoblasts, fetal fibroblasts, white blood cells (such as T lymphocytes, B lymphocytes and helper T cells), infected white blood cells, cells stem (eg, hematopoietic stem cells positive for CD34), epithelial cells, tumor cells and infectious organisms

(e.g., bacteria, protozoa and fungi).

Samples can be divided into multiple homogeneous components using the procedures described herein. Several similar devices containing different binding moieties specific for a series or parallel cell population can be connected. Serial separation can be used when seeking to isolate rare cells. On the other hand, parallel separation can be used when it is desired to obtain differential distribution of various populations in blood. Figures 20A and 20B show parallel and serial systems for the separation of various populations of blood cells. For parallel devices, two or more sets of obstacles that join different types of cells can be located in different regions or they can be intercalated with each other, e.g. eg, in a chessboard type pattern or in alternate rows. In addition, an obstacle game can be attached to the top of the device and another game can be attached to the bottom of the device. Each game can then be derivatized to unite different cell populations. Once a sample has passed through the device, the upper and lower upper can be separated to provide isolated samples of two different cell types.

The cell binding device can be used to remove the outflow of a given population of cells or to capture a specific population of cells expressing a specific surface molecule for further analysis. Obstacle-bound cells can be removed from the chamber for further analysis of the homogeneous cell population (Figure 21). This withdrawal can be achieved by incorporating one or more orthogonal inputs and outputs to the flow direction. The cells can be removed from the chamber at an increased flow rate, which has a greater shear stress, to overcome the bond strength between the cells and the obstacles. Other approaches may involve coupling junction moieties with reversible binding properties, e.g. eg, which are activated by pH, temperature or electric field. The binding moiety or the molecule bound on the surface of the cells can also be cleaved by enzymatic means or other chemical means.

In the example of isolation of fetal red blood cells, a sample that has passed through a lysis device is passed through a cell binding device, whose surfaces are coated with CD45. The white blood cells expressing CD45 present in the mixture bind to the walls of the device and the cells that pass through the device are enriched in fetal red blood cells. Alternatively, the obstacle surfaces and the device are coated with anti-CD71 in order to bind fetal nucleated red blood cells (which express the CD71 cell surface protein) of a whole maternal blood sample. One percent of adult white blood cells also express CD71. A sample of maternal blood is passed through the device and both populations of cells expressing CD71 bind to the device. This results in the removal of fetal red blood cells from the blood sample. Then, the fetal cells are collected and analyzed. For example, the cells are collected on a flat substrate for fluorescence in situ hybridization (FISH), followed by cell fixation and imaging. Figures 22A-22C show the use of FISH in a cell attached to an obstacle in a binding device of the invention. The cell, of fetal origin, is stained for X and Y chromosomes using fluorescent probes. These data show the feasibility of the optical image diagnosis of FISH stained cells in detection and diagnosis positions of chromosomal abnormalities.

Alternative realizations Another embodiment of the cell binding device uses chemically derivatized glass / plastic beads trapped in a poorly crosslinked hydrogel, such as, but not limited to, poly (vinyl alcohol), poly (hydroxy-ethyl methacrylate), polyacrylamide or polyethylene glycol (Figure 23 ). Chemically derivatized pearls serve as obstacles in this embodiment. A mixture of cells is directed into the cell removal device through two diametrically opposite entries. The positive pressure (e.g., of an infusion pump or liquid column) or negative pressure (e.g., of a syringe pump in traction mode, a vacuum pump or a vacuum cleaner) drives the liquid to through the hydrogel. The interaction of the cells in the sample with the chemically derivatized pearls dispersed in the three-dimensional volume of the hydrogel results in the removal of the cells, e.g. eg, white blood cells (negative selection), or cell capture, e.g. eg, fetal red blood cells (positive selection). Molecular weight, crosslinking density, pearl density and distribution and flow rates can be optimized to allow maximum interaction and capture of relevant cells by the beads. The high water content of the hydrogel provides a structure to trap the beads while allowing ease of flow through the sample. The sample is then collected through two diametrically opposed outlets. The branched inlet / outlet channel design guarantees maximum homogeneity in the distribution of the sample through the hydrogel volume.

In yet another embodiment, the beads are replaced by direct chemical derivatization of the hydrogel polymer side chains with the binding moiety (e.g., a synthetic ligand or monoclonal antibody (Acm)). This approach can provide a very high density of molecular capture sites and thus ensure a higher probability of capture. An added advantage of this approach is a potential use of the hydrogel-based cell removal device as a sensor for capturing fetal cells in the positive selection mode (select fetal cells with Acm), for example, if the chemistry of the structure The polymer and side chain is designed to capture fetal cells and also cross-link the hydrogel in the procedure. The cells bind to numerous side chains by antigen-Acm interaction and, therefore, serve as a crosslinker for the polymer chains and the reduction in flow output over time due to the increased crosslink density of the polymer can be matched. Mathematically the number of fetal cells captured in the 3D matrix of the polymer. When the desired number of fetal cells is captured (measured by reduction in outflow), the device can stop further processing the maternal sample and proceed to the analysis of fetal cells. Captured fetal cells can be released for analysis by using a photoactive coupling agent in the side chain. The photoreactive agent couples the target ligand or Acm to the polymer structure and, upon exposure to a pulse of UV or IR radiation, the ligands or Acm and the associated cells are released.

C. Cell selection

In this device, a mixture of cells from which unwanted cells have normally been removed is selected in a microfluidic device. An exemplary device for this stage is described in International Publication No. WO 01/35071. The cells of the device are then tested, e.g. eg, by microscopy or colorimetric assay, to locate the desired cells. The desired cells can then be analyzed in the set, e.g. eg, by lysis followed by PCR, or the cells can be collected from the set by a variety of mechanisms, e.g. eg optical tweezers. In the exemplary device described in WO 01/35071, the cells are introduced into the selection device and can be passively deposited in holes made in the device. Alternatively, positive or negative pressure can be used to direct the cells to the holes in the assembly. Once the cells have been deposited in the holes, the selected cells can be released individually from the set by activators, e.g. eg, bubble activated pumps. Other procedures to immobilize and release cells, e.g. eg, dielectrophoretic entrapment, can also be used in a selection device. Once released from the set, the cells can be collected and subjected to analysis. For example, a fetal red blood cell is identified in the set and then analyzed to assess genetic abnormalities. Fetal red blood cells can be identified morphologically or by a specific molecular marker (e.g., fetal hemoglobin, transferrin receptor (CD71), thrombospondin receptor (CD36) or glycophorin A (GPA)).

D. Separation based on size

Another device is a device for the separation of particles based on the use of sieves that allow the selective passage of particles according to their size, shape or deformability. The size, shape or deformability of the pores of the sieve determines the types of cells that can pass through the sieve. Two or more sieves can be placed in series or in parallel, e.g. eg, to remove cells of increasing size successively.

Device. In one embodiment, the screen includes a series of obstacles that are separated by a space. A variety of obstacle sizes, geometries and placements can be used in devices of the invention. Different forms of obstacles can be used in a sieve, e.g. eg, those with circular, square, rectangular, oval or triangular cross sections. The size of the gap between the obstacles and the shape of the obstacles can be optimized to ensure rapid and efficient filtration. For example, the size range of the GR is in the order of 5-8 µm, and the size range of the platelets is in the order of 1-3 µm. The size of all GB is larger than 10 µm. Large gaps between obstacles increase the speed at which GR and platelets pass through the sieve, but an increased hole size also increases the risk of losing WBC. Smaller hole sizes ensure a more effective capture of GB but also a slower pitch speed for GRs and platelets. Depending on the type of application, different geometries can be used.

In addition to obstacles, sieves can be manufactured by other procedures. For example, a sieve could be formed by molding, electroforming, etching, drilling or otherwise creating holes in a sheet of material, e.g. eg, silicon, nickel or PDMS. Alternatively, an inorganic matrix or polymer matrix (eg, zeolite or ceramic) having an appropriate pore size as a sieve could be employed in the devices described herein.

A problem associated with devices of the invention is the obstruction of the sieves. This problem can be reduced by appropriate sieve shapes and designs and also by treating sieves with non-adhesive coatings such as bovine serum albumin (BSA) or polyethylene glycol (PEG), as described herein. A procedure to avoid clogging is to minimize the contact area between the screen and the particles.

The scheme of a low shear filtration device is shown in Figure 24. The device has an inlet channel that leads into a diffuser, which is a widened portion of the channel. Normally, the canal widens with a V-shaped pattern. The diffuser contains two sieves that have pores with a shape to filter, for example, GR and smaller blood platelets, while enriching the population of fetal GB and GR. The diffuser geometry widens the laminar flow stream lines causing more cells to come in contact with the sieves as they move through the device. The device contains 3 outputs, two outputs collect cells that have passed through the sieves, e.g. eg, GR and platelets, and an output collects the enriched fetal GB and GR.

The diffuser device does not normally guarantee 100% removal of GR and platelets. However, the initial GR: GB ratios of 600: 1 can be improved at ratios around 1: 1. The advantages of this device are that the flow rates are low enough so that the shear stress on the cells does not affect the phenotype or viability of the cells and that the filters guarantee that all large cells (i.e. those that cannot pass through sieves) are retained so that the loss of large cells is minimized or eliminated. This property also guarantees that the population of cells that pass through the sieve does not contain large cells, although some smaller cells may be lost. Widening the diffuser angle will result in a larger enrichment factor. Major enrichment can also be obtained by serially placing more than one diffuser, where the output of a diffuser feeds the input of a second diffuser. The widening of the gaps between the obstacles can expedite the removal procedure with the risk of losing large cells through the larger pores of the sieves. To separate maternal red blood cells from fetal nucleated red blood cells, an exemplary spacing is 2 -4 µm.

Process. The device is a continuous flow cell separator, e.g. eg, that filters larger amounts of fetal GB and GR from blood. The location of the sieves in the device is chosen to ensure that the maximum number of particles come into contact with the sieves, while at the same time preventing blockage and allowing the recovery of the particles after separation. In general, the particles move through their laminar flow lines that are maintained due to an extremely low Reynolds number in the channels of the device, which are normally micron sized.

Manufacturing. Simple microfabrication techniques such as soft lithography with poly (dimethylsiloxane) (PDMS), polymer casting (e.g., using epoxies, acrylics or urethanes), injection molding, hot embossing of polymers, laser micromachining, micromachining can be used. of thin film surface, deep etching of glass and silicon, electroforming and 3D manufacturing techniques such as stereolithography for the manufacture of the channels and sieves of devices of the invention. Most of the procedures listed above use photomasks for duplication of micro features. For characteristics larger than 5 µm, transparency based emulsion masks can be used. The characteristics of sizes between 2 and 5 µm may require glass-based photomasks. For smaller features, a direct writing mask by glass-based electron beam can be used. Then, the masks are used to define a photoresist pattern for etching in the case of silicon or glass or to define negative replicas, e.g. eg, using SU-8 photoresist, which can then be used as a pattern for replica molding of polymeric materials such as PDMS, epoxies and acrylics. The fabricated channels can then be joined on a rigid substrate such as glass to complete the device. Other manufacturing processes are known in the art. A device of the invention can be manufactured from a single material or a combination of materials.

Example. In one example, a device for separation based on the size of GR and smaller platelets of the larger GBs was manufactured using simple smooth lithography techniques. A chrome photomask was made that had the characteristics and geometry of the device and was used to stamp a silicon wafer with a negative replica of the device in SU-8 photoresist. This pattern was then used to manufacture channel and sieve structures in PDMS using standard replica molding techniques. The PDMS device is attached to a glass sheet after plasma treatment with O2. Diffuser geometry is used to widen laminar flow stream lines to ensure that most of the particles or cells that flow through the device will interact with the sieves. The smaller GR and platelets pass through the sieves and the larger GBs are confined in the central channel.

Device combination

The devices described herein can be used alone or in any combination. In addition, the steps of the procedures described herein can be used in any order. A schematic representation of a combination device for detecting and isolating fetal red blood cells is shown in Figure 25. In one example, a sample can be processed using the cell lysis step and then the desired cells can be trapped with a binding device. cells. If the trapped cells are pure enough, no further processing step is required. Alternatively, only one of the lysis or binding steps can be used before selection. In another example, a mixture of cells can be subjected to lysis, separation based on size, binding and selection.

The methods of the invention can be carried out in an integrated device containing regions for cell lysis, cell binding, selection and size-based separation. Alternatively, the devices can be separated and the cell populations obtained at each stage can be collected and transferred manually to devices for additional processing steps.

Positive or negative pressure pumping can be used to transport cells through the microfluidic devices of the invention.

Analysis

After being enriched by one or more of the devices described herein, the cells can be collected and analyzed by various procedures, e.g. eg, nucleic acid analysis. The sample can also be processed further before analysis. In one example, the cells can be collected on a flat substrate for fluorescence in situ hybridization (FISH), followed by cell fixation and imaging. Such an analysis can be used to detect fetal abnormalities such as Down syndrome, Edward syndrome, Patau syndrome, Klinefelter syndrome, Turner syndrome, sickle cell anemia, Duchenne muscular dystrophy and cystic fibrosis. The analysis can also be performed to determine a specific trait of a fetus, e.g. eg sex.

Other realizations

Reference is made herein to all patent publications and patent applications mentioned in the previous specification in the context in which they are mentioned respectively. Various modifications and variations of the described system of the invention will be apparent to those skilled in the art without departing from the scope of the invention. Although the invention has been described in connection with specific embodiments, it should be understood that the invention as claimed should not be unduly limited to such specific embodiments. In fact, it is intended that various modifications of the ways described for carrying out the invention that are obvious to those skilled in the art are within the scope of the invention.

Other embodiments are in the claims.

Claims (12)

1. A microfluidic device comprising:

(to)

 a first region of fixed obstacles arranged in a microfluidic channel defining a liquid flow path, in which the obstacles of the first region preferably bind to a first type of cell compared to a second type of cell, in which the Obstacles are placed in at least two columns and at least two rows, in which the rows are placed perpendicular to the liquid flow path in relation to the obstacles in the previous row, thereby forming a set of equilateral triangular obstacles; Y

(b)

 a second region of fixed obstacles arranged in the microfluidic channel, in which the obstacles of the second region preferably join a third type of cell compared to a fourth type of cell, in which the obstacles are placed in at least two columns and at least two rows, in which the rows are placed perpendicular to the liquid flow path and the obstacles of each successive row are displaced in a direction perpendicular to the liquid flow path in relation to the obstacles of the previous row , thus forming a set of equilateral triangular obstacles,

in which the second region is located beyond the first region in the microfluidic channel.

2.

The microfluidic device of claim 1, wherein the obstacles are coated with an antibody.

3.

The microfluidic device of claim 1 or 2, wherein an obstacle spacing is at least 50 µm.

Four.

The microfluidic device of any one of claims 1 to 3, wherein a spacing between obstacles is 100 µm maximum.

5.

The microfluidic device of any one of claims 1 to 4, wherein the obstacles of the first region, the second region, or both first and second regions, are substantially uniform in size.

6.

The microfluidic device of any one of claims 1 to 5, further comprising a pump.

7.

The microfluidic device of any one of claims 1 to 6, wherein the device is optically transparent or has transparent windows.

8.

The microfluidic device of any one of claims 1 to 7, wherein the obstacles are in contact with both the top and bottom of the chamber.

9.

The microfluidic device of any one of claims 1 to 8, wherein the obstacles have a cylindrical cross section.

10.

The microfluidic device of any one of claims 1 to 9, wherein the device is made from a polymer.

eleven.

The microfluidic device of any one of claims 1 to 10, wherein the obstacles are positioned to allow the flow of cells without being mechanically compressed between the obstacles and, therefore, damaged, during the flow process.

12.

The use of a device according to any of claims 1 to 11 for medical diagnosis.